Customized load-bearing and bioactive functionally-graded implant for treatment of osteonecrosis

ABSTRACT

An engineered medical device for treatment of osteonecrosis is provided where the size, porosity and ceramic content of the device can be personalized based on an individual patient&#39;s anatomical and physiological condition. The device distinguishes different segments mimicking anatomically-relevant cortical and cancellous segments, in which the cortical segments of the device can sustain mechanical loading, and the cancellous segment of the device can promote bone ingrowth, osteogenesis and angiogenesis.

FIELD OF THE INVENTION

This invention relates to devices and methods for preventing progressionof osteonecrosis at its early stage. In particular, the inventionrelates to an implant for reconstruction of the osteonecrotic area.

BACKGROUND OF THE INVENTION

Osteonecrosis of the hip (ONH) is a debilitating disease that isincreasing in incidence worldwide and frequently progresses to collapseof the femoral head and osteoarthritis that necessitates total hipreplacement. In the early stages of ONH, various medical and surgicaltreatments to preserve the integrity of the femoral head includingpharmacological agents such as statins and anticoagulants,electromagnetic therapy, weight reduction, protected weight bearing andrange of motion exercises have been attempted with limited success andhave not prevented collapse or provided lasting improvement.Alternatively, surgical treatments such as core decompression of thenecrotic segment are often performed to relieve pain prevent progressionat the early stages of ONH, prior to femoral head collapse. However awide range of the success rates have been reported from 20-70% for theearly (pre-collapse) stages. Maintenance of sphericity of the patient'sown femoral head requires both mechanical and biological strategies towithstand intermittent loading and, at the same time, reconstitute thenecrotic femoral bone segment. To prevent the collapse of the femoralhead after core decompression, the removed necrotic tissue could befilled by a graft or implant to facilitate reconstruction of thenecrotic area.

A vascularized fibula graft is one current clinical option, but itsuffers from several limitations including pain associated with graftharvesting, availability of sufficient transplantable bone, and thepossibility of infection as well as donor site morbidity. Currently, theimplant manufacturer Zimmer supplies a porous tantalum metal implant asa treatment for early stage ONH. However, the tantalum metal implantwill remain in the proximal femur region for the patient's lifespan.Furthermore, studies have reported no evidence of vascular invasion andminimal bone ingrowth (only 1.9%) in such implants, much less than themean density (26.2%) of adjacent femoral head trabecular bone. Acontinuous shell of new cortical bone forms around the tantalum implantthat blocks vascular and cancellous invasion. This nullifies the purposeof core decompression and leads to a gradual increase in theintramedullary pressure with subsequent pain. Progressive severe painleads to clinical failure necessitating total hip replacement whereasthe presence of the tantalum metallic implant complicates subsequentsurgical procedures including total hip replacement.

SUMMARY OF THE INVENTION

The present invention advances the art and provides technology fortreating early stage ONH when the femoral head is still round andtherefore salvageable. After removing the necrotic tissue, amechanically- and biologically-sound biodegradable template is used forreconstitution of the osteonecrotic area within the femoral head.

In this invention, we developed a mechanically robust andfunctionally-graded scaffold (FGS) with spatiotemporally-controlleddegradation and mechanical to properties as a filler for the coredecompression tunnel. The FGS for ONH treatment was composed of threesections/segments of spatially-graded porosity: a proximal segment tosupport the subchondral area, a middle (intermediary) section/segmentlocated in the main osteonecrotic area of the femoral head, and a distalsection/segment to support the cortical-like structure (FIG. 1).

The proximal and distal segments of the scaffold were designedrelatively less porous to mimic cortical bone, and provide support towithstand compressive mechanical loads and maintain the integrity of thearticular surface. The middle segment was designated to possessrelatively higher porosity to mimic the trabecular bone of the femoralhead, which provides an appropriate template for vascularization and newbone ingrowth and mechanical stability against bending moments appliedto the femoral head. In one embodiment, we utilized additivemanufacturing (AM) technology to form FGSs with controlled porosity forvarious segments of the FGSs in a single-stop-shopping process. AMtechnology also enables the customization of FGS for individual patientsbased on the anatomical size and level of ONH condition. Biodegradablepoly(ε-caprolactone) (PCL) and β-tricalcium phosphate (β-TCP) were usedfor the fabrication of FGS because PCL and β-TCP are in current clinicaluse and can provide an appropriate scaffold for bone tissue engineering.We first characterized the physical and mechanical properties of eachsegments of FGS. A comprehensive animal study was conducted to show thepotential of design and material of FGS for promoting bone ingrowth andregeneration in core decompression bone tunnel of rabbit femoral headsto mimic treatment of ONH at early stages.

In one embodiment, an engineered medical device for treatment ofosteonecrosis is provided. The device is a cylindrically-shapedbiodegradable scaffold made of filaments distinguishing three sectionsof spatially graded chemical composition, porosity and mechanicalstrength. The three sections define (1) a proximal section, (2) a distalsection and (3) an intermediary section in between the proximal anddistal sections. The intermediary section is longer than each of theproximal and distal sections. The proximal and distal sections have aporosity that is less than the porosity of the intermediary section. Theproximal and distal sections have a mechanical strength that is higherthan the mechanical strength of the intermediary section.

In one example, the proximal section has a porosity that is less thanthe porosity of the distal section.

In another example, the proximal section has a mechanical strength thatis less than the mechanical strength of the distal section.

In still another example, the scaffold is a made of a polymer and aceramic. The scaffold is made of polycaprolactone (PCL), calciumphosphate, beta-tricalcium phosphate (beta-TCP), hydroxyapatite, or acombination thereof. The three sections could include calcium phosphateor beta-tricalcium phosphate and with that the proximal and distalsections could have a higher calcium phosphate or beta-tricalciumphosphate concentration than the intermediary section resulting in themechanical strength difference, an osteoconductivity difference, and/ora degradation difference between proximal and distal sections comparedto the intermediary section.

In still another example, the mechanical strength difference resultsfrom the difference in the porosity between proximal and distal sectionscompared to the intermediary section.

In still another example, the mechanical strength for the sectionsvaries in a range of 0.5 to 6 MPa, with a stiffness for the sectionsvarying in a range of 20 to 100 Mpa, while preserving that themechanical strength of the proximal and distal sections is higher thanthe mechanical strength of the intermediary section.

In still another example, the porosity difference changes gradually fromthe proximal section to the intermediary section and to the distalsection, while preserving that the porosity of the proximal and distalsections is less than the porosity of the intermediary section.

In still another example, the mechanical strength difference changesgradually from the proximal section to the intermediary section and tothe distal section, while preserving that the mechanical strength of theproximal and distal sections is higher than the mechanical strength ofthe intermediary section. In still another example, the porosity of theproximal and distal section is defined between 0% to 40% and wherein theporosity of the intermediary section is defined above 40% to 95%.Alternatively, the porosity of the proximal and distal section isdefined between 0% to 30% and wherein the porosity of the intermediarysection is defined above 30% to 95%.

In still another example, the scaffold has a biodegradation rate thatmatches bone regeneration, wherein the biodegradation rate can betailored specifically to a patient's needs. Alternatively, the threesections each have a biodegradation rate that matches bone regenerationintended and designed for its respective sections.

In still another example, the medical device is adapted to fit in atunnel bored in a femoral head intended for reconstruction of anosteonecrotic area.

In still another example, one or more growth factors or cellularcomponents are added to the surface of the scaffold.

In still another example, the scaffold is a three-dimensionally printedscaffold. In still another example, the proximal and distal sectionsmimic cortical bone. In another example, the proximal section mimics asubcondoral surface profile, and/or in another example, the intermediarysection is intended to replace trabecular bone of necrotic tissue. Inone example, the intermediary section is a template for vascularizationand bone ingrowth.

In yet another example, the size, porosity and ceramic content of thedevice is personalized based on an individual patient's anatomical andphysiological condition. A skilled artisan could envision embodiments ofa scaffold including more than three sections or segments based on thesame approach and philosophy as discussed herein for a scaffold withthree segments.

The FGS offers advantages over existing technologies. First, such asynthetic graft prototype of FGS seamlessly integrates both mechanicaland biological strategies by mimicry of anatomically-relevant corticaland cancellous segments, in which the cortical segments of the FGS cansustain mechanical loading, and the cancellous segment of the FGS canpromote bone ingrowth, osteogenesis and angiogenesis. In addition, theopen pore structure and bioresorbable nature of our implants will havean improved integration capability compared to fibular autograft, andporous non-resorbable tantalum implants because both of the currentlyavailable grafts either possess or generate an impermeable cortical bonyshell to impair vascular invasion and osseointegration. Second, thecombination of the 3D printing and the novel biodegradable biomaterialsthat are suitable for 3D printing allows us to personalize the implantsregarding size, anatomical shape, and more uniquely, properties such astempo-spatial degradation profiles to better facilitate revitalizationof the osteonecrotic area of ONH. Third, the FGS is made ofpolymer-ceramic composite and can be customized in the operating room bycutting and trimming with surgical instruments to fit the patient'sspecific defect. This feature renders available to patients eithercatalog-based off-of-shelf products or customized products. Fourth, thepersonalizable biodegradable FGS can be readily combined with theclinically available cell therapy to further improve the clinicaloutcome.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 shows a schematic of the functionally-graded scaffold (FGS)according to an exemplary embodiment of the invention.

FIG. 2 shows digital photos of a representative FGS according to anexemplary embodiment of the invention.

FIG. 3 shows a representative micro CT image of a longitudinalcross-section of FGS according to an exemplary embodiment of theinvention.

FIG. 4 shows SEM image of scaffold strut surface according to anexemplary embodiment of the invention.

FIG. 5 shows SEM image of scaffold strut surface after NaOH treatmentaccording to an exemplary embodiment of the invention.

FIGS. 6A-C show physical characterization of a scaffold according to anexemplary embodiment of the invention.

FIGS. 7A-D show microCT images of rabbit femoral heads in the presenceand absence of FGS at 8 weeks after core decompression according to anexemplary embodiment of the invention.

FIGS. 8A-C show quantitative analysis of scaffold degradation and boneingrowth in the FGS at 8 weeks after implantation according to anexemplary embodiment of the invention.

FIGS. 9A-C show bone ingrowth in the FGS in the non-decalcifiedhistology sample according to an exemplary embodiment of the invention.

FIGS. 10A-C show bone ingrowth and regeneration in the decalcifiedhistology samples according to an exemplary embodiment of the invention.

FIGS. 11A-D show TRAP positive cells in the FGS according to anexemplary embodiment of the invention.

DETAILED DESCRIPTION

Materials and Methods

Medical-grade poly(ε-caprolactone) (mPCL) pellets (Aldrich ChemicalCompany) with density of 1.145 g/cm³ (Mn=80,000) were used as received.β-tricalcium phosphate (β-TCP) powder with a specific surface area of 17m²/g and particle size of mesh 60 was obtained from Nanocerox, Inc. (AnnArbor, Mich.). A slurry was prepared by dispersing β-TCP powder inacetone at 50° C., followed by gradual incorporation of PCL pellets. APCL:β-TCP:acetone weight ratio of 80:20:400 was used. Vigorousmechanical stirring was performed for 3 h to ensure the homogenousdistribution of the β-TCP particles in the PCL slurry. The slurry waspoured out evenly into glass plates, left overnight to evaporate, andfurther dried in a vacuum overnight to remove any residual acetone.Then, as raw material for AM process, the material was shaped as longfilaments via warming up and extruding through a 2.2-mm orifice of anin-house-made filament-maker machine.

Design and Fabrication of PCL-β-TCP FGS

FGS for ONH treatment was designed to be composed of three segments ofspatially-graded porosity, including 3.5 mm in diameter×4 mm in lengthfor proximal segment of 15% porosity, 3.5 mm in diameter×17 mm in lengthfor middle segment and 3.5 mm in diameter×6 mm for distal segment ofFGS. Such a graded scaffold was formed in one piece using ourin-house-built layer-by-layer 3D fabrication system, Hybprinter.Briefly, the filaments were fed into the machine AM machine where theywere melted at 140° C., extruded as 350 μm struts and laid down in 0/90°patterns layer-upon-layer to form porous lattice-shaped scaffolds. Thelayer thickness (=distance between the sequential layers) was selectedequal to 300 μm. The spacing between the adjacent struts was defined ina way to provide 15% porosity in proximal segment and distal regions and60% porosity in middle region. Also, similar setting was used to formsamples for physical and mechanical characterizations. All scaffoldswere surface-treated in 5M NaOH for 12 h at room temperature to enhancetheir hydrophilicity and micro-roughness on the strut surface.

Physical Characterizations

Porosity

Porosity of samples for proximal, middle, and distal segments wasmeasured using high-resolution micro-CT scanner (SMX-100CT-SV3; ShimadzuCo., Kyoto, Japan). The entire set of radiographs was deconvoluted bycomputer software to reconstruct a 3D image of the microstructure with avoxel size of 12 μm³. The 3D data were processed with commerciallyavailable 3D image processing software (VG Studio MAX 2.0; VolumeGraphics, Heidelberg, Germany), and the porosity of each segment of thescaffold was measured from the binary material images. The spatialboundary between the pores and the scaffold was determined based on theremarkable difference of their intensities.

Scanning Electron Microscopy

AM-made scaffolds were also scanned via secondary electron emissionimaging using SEM (FEI XL30 Sirion) to analyze strut size, pore size,and layer thickness as well as to observe surface morphology of thestruts. Three samples per group were used. The PCL-β-TCP scaffolds werefirst sputter-coated with gold (10 nm) (SPI Sputter, SPI SupplierDivision of Structure Prob Inc., West Chester, Pa., USA) to make themelectrically conductive. The SEM acceleration voltage was set to 5 kV.

Water-Uptake

The water-uptake characteristic represents hydrophilicity of thescaffolds and the surface condition of the scaffold struts for cellanchorage, adhesion and proliferation as well as new tissue ingrowth andintegration. The water uptake ability of the porous scaffolds wasdetermined by completely immersing the samples in distilled water for 60s. The bulk water accumulated on the outer surfaces of the samples wasremoved via blotting with a piece of wet filter paper. Then, the wetscaffolds were weighed immediately using an electronic balance (XS105,METTLER TOLEDO, Columbus, Ohio). The percentage of water uptake wasdetermined as

${{WaterUptake}\mspace{14mu}\%} = {\frac{W_{w} - W_{d}}{W_{d}} \times 100}$where W_(w) and W_(d) are the weight of wet and dry samples,respectively. The test was conducted using five specimens of each sampletype (i.e., different strut size). Each measurement was triplicated foreach scaffold and the average value was calculated. Five samples pergroup were tested.Degradation

The degradation of PCL-β-TCP scaffolds with 15% and 60% porosity wascompared in alkaline medium. To accelerate the hydrolysis reaction 5MSodium hydroxide (NaOH) was used. Scaffolds of 5×5×10 mm³ were submergedin 2 mL of NaOH at 37 degrees Celsius. The PCL scaffolds were dried andweighted after 12 hr, 24 hr, 36 hr and 48 hr. The degradation rate wasdetermined as weight loss percentage at each time point.

Mechanical Characterization

Compression Test

The mechanical properties of the scaffolds were measured using anInstron 5944 uniaxial testing system with a 2 kN load-cell (InstronCorporation, Norwood, Mass.). Specimens with 5×5 mm² squarecross-section and 10 mm in height were mechanically tested undercompression between two platens (one self-aligning, one fixed) at aspeed of 0.1% strain/sec up to 25% strain with 1 N preload. Displacementwas determined from an extensometer (Epsilon Technology Corp, JacksonWyo.) attached to the two platens. Five samples were tested for eachgroup of 15% and 60% porosity. The testing protocol was adapted from thereference (22) where, for all the specimens, the apparent modulus ofelasticity was calculated as the slope of the initial linear portion ofthe stress-strain curve. The effective stress values were determined asthe compressive loading value per the apparent initial cross-sectionalarea of each test specimen. The strain values were calculated viadividing the deformation values with the initial specimen height. Yieldcompression strength was defined as the intersection of thestress-strain curve with a line with a slope equal to the modulusstarting at an offset of 1% strain. Five samples per group weremeasured.

Flexural Test

To measure bending flexural stiffness of scaffolds, 3.7 mm×3.9 mm×70 mmsamples of 15% and 60% porosity were tested under 3-point bendingfollowing ASTM D790 instructions. Bending tests were conducted using asupport span of 56 mm, which resulted in a support span to specimendepth ratio of 16:1 (depth=3.5 mm). A 100 N load sensor was used for thetest. Following application of a 0.1 N preload, the central loading nosewas lowered at a rate of 0.1% strain/sec to 5% strain. The apparentflexural modulus was calculated as the slope of the initial linearportion of the stress-strain curve (between 0.1 and 1.1% strain). Fivesamples per group were measured.

In Vivo Evaluation

Animals

Five male New Zealand White rabbits (Charles River Laboratories Inc.USA) weighing from 3.5 to 4.0 kg were used in this study. For operativeprocedure, animals were anesthetized by administration of 40 mg/kgketamine and 4 mg/kg xylazine. Additional inhalation anesthesia wasperformed with isoflurane. Analgesia was administered by injection of0.5 mg/kg buprenorphine twice a day for the first 2 days after surgery.All experiments with animals were performed following StanfordUniversity Animal Care and Use Committee guidelines. All researchanimals has been approved by Stanford APLAC #28999, following approvedguidelines by the Stanford University's Institutional Review Board.

Surgical Procedure

A lateral skin incision was created to expose the greater trochanter.Core decompression of 3.5 mm diameter was performed from distal end ofthe greater trochanter along the axis of the femoral neck on both hips.The tunnel direction was in the mid axis of the femoral neck. First, abone tunnel was created from laterally towards the femoral headsuperomedially using a 2 mm diameter drill under fluoroscopic guidance.The depth of bone tunnel was approximately 27 mm. This 2 mm tunnel waswidened with a 3.5 mm diameter drill bit. FGS (3.5 mm in diameter and 27mm in length) was randomly assigned and inserted into the tunnel in theright or the left hip and the other tunnel was kept empty as a control.The scaffold was press fitted into the bone tunnel. Five implants wereused for this study based on our previous experience where bone tunnelwas created on the rabbit femur. Power analysis indicated that a samplesize of five per group would provide 80% statistical power to detectsignificant differences between the groups (α=0.05, β=0.20) usinganalysis of variance (ANOVA).

Micro-CT Analysis

A microfocus X-ray computed tomography system (SMX-100CT-SV3; ShimadzuCo., Kyoto, Japan) was used to acquire microstructural information forproximal femur at 8 weeks after implantation of FGSs. The entire set ofradiographs was deconvoluted by computer software to reconstruct a 3Dimage of the microstructure with a voxel size of 12 μm and was evaluatedusing 3D image-processing software VG studio MAX 2.0 (Volume Graphics,Heidelberg, Germany). A cylindrical region of interest (3.5 mm indiameter×4 mm in length for proximal segment, 17 mm in length for middlesegment and 6 mm for distal segment of FGS) was co-centricallypositioned over the core decompression site. The location of originalbone tunnel in empty-tunnel group was confirmed on CT image based on toanterior-posterior and lateral view of X-ray taken during surgery. Thevolume of new bone and remaining FGS was determined by the software.Thresholds were applied to differentiate between new bone and residualscaffold material in the region of interest.

Histology

After 8 weeks of implantation, following the euthanasia, the proximalfemoral bone was harvested and prepared for histology. Three samplesfrom each group were prepared for decalcified histological analysis. Thespecimens were fixed in 10% phosphate-buffered formaldehyde (pH 7.25)for 24 h and decalcified in 15% EDTA (pH 8.0) at 4° C. Completedecalcification was confirmed by X-ray. Then, the samples weredehydrated in graded ethanol (70, 85, 90 and 100%), then cleared inxylene, and embedded in paraffin. Thin sections (8 μm) were cut andmounted on glass slides. Before staining, the sections weredeparaffinized in 100% xylene and rehydrated in graded ethanol.Specimens were analyzed by hematoxylin and eosin (H&E) staining andtartrate resistant acid phosphate (TRAP) staining. The number of TRAPpositive multinucleated cell was counted. Two samples from each groupwere prepared for undecalcified histology. The specimens were dehydratedin graded ethanol after fixed in 10% formaldehyde, and then embedded ina cold setting epoxy resin. Thick specimens (250 μm) were cut with aband saw, and ground to a thickness of 50 μm. Each section was evaluatedwith Stevenel's blue and Van Gieson's picrofuchsin staining. TRAPstaining was also performed.

Statistical Analysis

All data are expressed as means±standard deviation (SD). For thecomparison among three segments (proximal, middle and distal), thehomogeneity of the variance among groups was assessed with the Bartletttest before examined with ANOVA. When the variance was homogeneous,comparisons between groups were performed with one-way ANOVA followed bya post hoc test (Tukey-Kramer multiple comparison test). All analyseswere performed using JMP 9 (SAS Institute, Cary, N.C.). For comparisonbetween two groups (empty group vs. scaffold group and low porositysegment vs. high porosity segment), Student's t test was used toinvestigate the significant difference between the groups. Values ofp<0.05 were considered statistically significant.

Results

Morphology of FGS

FIG. 2 show digital photos of a representative FGS. The proximal anddistal less porous segments are shown from both top and side views. FIG.3 shows a representative micro CT image of a longitudinal cross-sectionof FGS. These images confirmed the successful fabrication of a FGS withtwo less porous segments in the two ends and a more porous middlesegment. Also, the SEM image of scaffold strut surface is shown in FIGS.4 and 5, demonstrating roughness (FIG. 4) and micro-pores (FIG. 5)formed after NaOH treatment.

Porosity and Water-Uptake

FIG. 6A shows the porosity and water-uptake of scaffolds for high andlow porosity segments of FGS construct. The porosities of scaffolds viamicro CT analysis were 16.8±1.8%, 59.5±1.2% and 16.4±1.7% for proximal,middle and end segment, respectively, which were in good agreement(˜2.0% difference) with designed porosity. Water-uptake yielded valuesof 15.1±2.1% for dense segments and 72.3±12.9% for the less densesegment, respectively. There was no significant difference inwater-uptake and porosity measurements, demonstrating excellenthydrophilicity of sample surface treated by alkali.

Degradation

FIG. 6B shows the accelerated degradation rates of scaffolds of high andlow porosity in alkaline medium (5M NaOH) for every 12 hr up to 48 hr.Both scaffolds exhibited almost linear degradation rates, and thescaffolds of 60% porosity demonstrated significantly higher weight lossrate in vitro compared to those of lower porosity under such anaccelerated degradation condition.

Mechanical Properties

The strength and apparent modulus of elasticity of PCL-β-TCP scaffoldswith low and high porosity were measured under compression. As shown inFIG. 2C, the strength of the 60% porous scaffold was 2.2 MPa in averagewhich is comparable with the strength of human trabecular bone (rangingfrom 0.2 to 10 MPa). The apparent modulus of 60% porous scaffolds wasabout 51.5 MPa. The scaffolds with lower porosity exhibitedsignificantly higher compressive strength and apparent elastic modulus:9.5 MPa and 213.4 MPa, respectively. Also, flexural mechanical stiffnessof the 60% porous scaffolds under three point bending condition wasmeasured equal to 104.8±19.7 MPa.

Gross Inspections

All rabbits tolerated the operation well. No infection of the operationsite or migration of the implant was seen in dissection aftereuthanasia. No apparent adverse reactions including inflammation orforeign body reactions were observed around implants.

Micro CT Analysis

Micro CT was used to measure the volumes of residue FGS and newly formedbone of explants at 8 weeks after surgery. FIGS. 7A-D showrepresentative micro CT images of explants with and without FGS. The CTimages showed that the lattice pattern and structural integrity of thescaffold were well maintained at 8 weeks after implantation. In the FGSgroup, the mineralized tissue was observed in the center of FGS infemoral head (FIG. 7B). In the empty-tunnel group, most of the space inthe bone tunnel remained empty, but there was some mineralized tissuearound the tip of the tunnel (FIG. 7D).

FIG. 8A shows the degradation rate of the FGS that was quantified usingCT analysis. The degradation rates of proximal, middle and distalsegments were 25.7±11.3%, 11.1±4.0%, and 20.4±6.9%, respectively. Thedifference in degradation rate between proximal and middle segments wasstatistically significant (p=0.028), whereas the differences indegradation rate between proximal and distal segments, or between middleand distal segments were not (p=0.539 and 0.180, respectively).

The volumes of newly formed bone measured at 8 weeks post surgery weremeasured 11.9±3.6 mm³ in the proximal segment, 8.2±5.8 mm³ in the middlesegment, 10.9±4.6 mm³ in the distal segment for the scaffold group, andwere 11.5±4.9 mm³ for the proximal segment, 4.5±2.6 mm³ for the middlesegment and 11.7±2.8 mm³ for the distal segment for the empty group(FIG. 8B). The bone ingrowth ratio for the scaffold group was 80.1±12.1%for the proximal segment, 7.7±5.2% for the middle segment and 57.7±25.4%for the distal segment (FIG. 8C). The bone ingrowth ratio for theempty-tunnel group was 30.5±13.5% for the proximal segment, 2.8±1.6% forthe middle segment and 20.5±4.8% for the distal segment. The differencein bone ingrowth ratio between the scaffold-filled group and theempty-tunnel group was statistically significant for the proximalsegment (p=0.0053) and distal segment (p=0.031), but not for the middlesegment (p=0.164).

Histology

At 8 weeks after implantation, considerable amount of new bone was seenin the undecalcified histology samples of both the FGS group (FIG. 9A)and empty-tunnel groups (FIG. 9B). In the FGS group, the newly formedbone tissue in the porous area of FGS was in direct contact with thescaffold strut (FIG. 9C). The images from paraffin embedded decalcifiedsamples showed that almost all space available was occupied by newlyformed bone in proximal segment of FGS (FIG. 10A). At high magnificationimages, Haversian canal- and Volkmann's canal-like structures wereobserved containing blood cells even near the center of the FGS (FIG.10C). The number of TRAP positive cell on the surface of the scaffoldwas 5.5±1.7 per mm² in the proximal segment, 3.8±1.8 per mm² in themiddle segment, and 7.1±2.2 per mm² in the distal segment. Thedifference in the TRAP positive cell number between the distal segmentand middle segments was significant, while the difference betweenproximal segment and middle segment, and proximal segment and distalsegment was not statistically significant (FIGS. 11A-D).

What is claimed is:
 1. An engineered medical osteonecrosis treatmentdevice, comprising: a cylindrically-shaped biodegradable scaffold madeof filaments having three cylindrical sections of spatially gradedchemical composition, porosity and mechanical strength, wherein thethree cylindrical sections distinguish in cylindrical alignment witheach other a proximal cylindrical section, a distal cylindrical sectionand a single intermediary cylindrical section positioned in between theproximal and distal cylindrical sections, wherein the intermediarycylindrical section is longer than each of the proximal and distalcylindrical sections, wherein a porosity of the proximal and distalcylindrical sections across the entire proximal and distal cylindricalsections is less than a porosity across the entire intermediarycylindrical section, and wherein a mechanical strength across the entireproximal and distal cylindrical sections is higher than a mechanicalstrength across the entire intermediary cylindrical section.
 2. Theengineered medical osteonecrosis treatment device as set forth in claim1, wherein the proximal cylindrical section has a porosity which is lessthan the porosity of the distal cylindrical section.
 3. The engineeredmedical osteonecrosis treatment device as set forth in claim 1, whereinthe proximal cylindrical section has a mechanical strength which is lessthan the mechanical strength of the distal cylindrical section.
 4. Theengineered medical osteonecrosis treatment device as set forth in claim1, wherein the scaffold is a made of a polymer and a ceramic.
 5. Theengineered medical osteonecrosis treatment device as set forth in claim1, wherein the scaffold is made of polycaprolactone (PCL), calciumphosphate, beta-tricalcium phosphate (beta-TCP), hydroxyapatite, or acombination thereof.
 6. The engineered medical osteonecrosis treatmentdevice as set forth in claim 1, wherein the three cylindrical sectionscomprise calcium phosphate or beta-tricalcium phosphate and wherein theproximal and distal cylindrical sections have a higher calcium phosphateor beta-tricalcium phosphate concentration than the intermediarycylindrical section resulting in (i) the mechanical strength difference,(ii) an osteoconductivity difference, (iii) a degradation differencebetween proximal and distal cylindrical sections compared to theintermediary cylindrical section, or a combination thereof.
 7. Theengineered medical osteonecrosis treatment device as set forth in claim1, wherein the mechanical strength difference results from thedifference in the porosity between proximal and distal cylindricalsections compared to the intermediary cylindrical section.
 8. Theengineered medical osteonecrosis treatment device as set forth in claim1, wherein the mechanical strength for the cylindrical sections variesin a range of 0.5 to 6 MPa, wherein a stiffness for the cylindricalsections varies in a range of 20 to 100 Mpa, while preserving that themechanical strength of the proximal and distal cylindrical sections ishigher than the mechanical strength of the intermediary cylindricalsection.
 9. The engineered medical osteonecrosis treatment device as setforth in claim 1, wherein the porosity difference changes gradually fromthe proximal cylindrical section to the intermediary cylindrical sectionand to the distal cylindrical section, while preserving that theporosity of the proximal and distal cylindrical sections is less thanthe porosity of the intermediary cylindrical section.
 10. The engineeredmedical osteonecrosis treatment device as set forth in claim 1, whereinthe mechanical strength difference changes gradually from the proximalcylindrical section to the intermediary cylindrical section and to thedistal cylindrical section, while preserving that the mechanicalstrength of the proximal and distal cylindrical sections is higher thanthe mechanical strength of the intermediary cylindrical section.
 11. Theengineered medical osteonecrosis treatment device as set forth in claim1, wherein the porosity of the proximal and distal cylindrical sectionis defined between 0% to 40% and wherein the porosity of theintermediary cylindrical section is defined above 40% to 95%.
 12. Theengineered medical osteonecrosis treatment device as set forth in claim1, wherein the porosity of the proximal and distal cylindrical sectionis defined between 0% to 30% and wherein the porosity of theintermediary cylindrical section is defined above 30% to 95%.
 13. Theengineered medical osteonecrosis treatment device as set forth in claim1, wherein the scaffold has a biodegradation rate that matches boneregeneration, wherein the biodegradation rate can be tailoredspecifically to a patient's needs.
 14. The engineered medicalosteonecrosis treatment device as set forth in claim 1, wherein thethree cylindrical sections each have a biodegradation rate that matchesbone regeneration intended and designed for its respective cylindricalsections.
 15. The engineered medical osteonecrosis treatment device asset forth in claim 1, wherein the medical device is adapted to fit in atunnel bored in a femoral head intended for reconstruction of anosteonecrotic area.
 16. The engineered medical osteonecrosis treatmentdevice as set forth in claim 1, wherein one or more growth factors orcellular components are added to a surface of the scaffold.
 17. Theengineered medical osteonecrosis treatment device as set forth in claim1, wherein the scaffold is a three-dimensionally printed scaffold. 18.The engineered medical osteonecrosis treatment device as set forth inclaim 1, wherein the proximal and distal cylindrical sections mimiccortical bone.
 19. The engineered medical osteonecrosis treatment deviceas set forth in claim 1, wherein the proximal cylindrical section mimicsa subchondral surface profile.
 20. The engineered medical osteonecrosistreatment device as set forth in claim 1, wherein the intermediarycylindrical section is intended to replace trabecular bone of necrotictissue.
 21. The engineered medical osteonecrosis treatment device as setforth in claim 1, wherein the intermediary cylindrical section is atemplate for vascularization and bone ingrowth.
 22. The engineeredmedical osteonecrosis treatment device as set forth in claim 1, whereinsize, porosity and a ceramic content of the device is personalized basedon an individual patient's anatomical and physiological condition.